The present disclosure concerns an apparatus and method for compensating for variations in the overall transfer ratio from transmitter input to receiver output between two or more sensors mounted at different anatomical locations. This includes variations due to overall changes in path length arising from different equivalent transmission characteristics through different body locations. Though also suitable for many other skin contact sensors using non-optical transduction principles, such as RF, magnetic field, electric field or vibration, the apparatus and method will be illustrated below by the example of an infrared (IR) wavelength transducer combining receiver and transmitter functions, which can be attached to the lower limbs of a patient.
Physically, a typical non-invasive optical transducer will consist of a cylindrical body with a slightly larger circular base. The base contains two small circular apertures within which are the tips of infrared transmitter and receiver devices. These are spaced apart with a centre-to-centre distance of a few mm.
For correct operation, the transducer base needs to be in accurate contact with the skin of a patient. This enables a beam of infrared energy to enter the skin and surface tissues to a depth in excess of 2 mm. Random photon to photon scattering within the beam enables a much attenuated signal to be gathered by the receiver device. The transmission path of the photons within the received signal has passed through both the capillary loops and subsurface horizontal structure of small veins and arteries which lies around 2 mm below the skin in a normal healthy adult human.
The overall transmission attenuation from transmitter to receiver is composed of several elements, including back-scatter and absorption by various anatomical structures such as tissue, ligament etc. That part of the transmitted intensity which reaches sub-surface blood vessels is further reduced by blood volume related absorption and therefore by the time-dependent quantity of blood in both arterial and venous elements of the vasculature.
However the physical spacing of the transmitter and receiver devices is such that the received signal is dominated by the blood volumes in the micro-circulation, rather than the larger and deeper veins and arteries. Processing of this received micro-circulation dependent signal in the time and frequency domains can reveal many valuable items of information, which correlate with diseases of both the vasculature and the peripheral nervous system.
There are several components to the IR energy exiting the tissue which is picked up in the receiver device of the transducer, as follows:    (a) absorption by pulsating and non-pulsating arterial blood, typically 1% to 2% of the total;    (b) absorption by venous blood and other tissue, typically <20% of the total; and    (c) scattering and reflection of incident light, typically >80% of the total.
In pulse oximetry, only the signal element dependent on pulsating arterial blood is important and the first processing step is to normalise this by dividing by the “dc” component of the received signal which is composed of all the other elements listed above. Averaging a period of received signal or passing the complete signal through a low-pass filter set to remove frequencies above, say, 0.05 Hz will produce a signal value representing the overall effect of scattering and reflection, tissue absorption, non-pulsatile arterial blood absorption, non-time dependent venous blood absorption and averaged low frequency vasomotor time dependent venous blood absorption.
Normalising by this “dc” signal provides a result which is unaffected by fluctuations in the incident light intensity and is effectively self-calibrating. Thus variations in scattering and reflection of incident light on entering the skin, as well as the transduction ratios of transmitter and receiver and circuit gain tolerances, do not affect the result.
In the alternative situation where the received signal component of interest comes from the time varying element of the venous blood, this signal is composed of:    (a) a breathing component with energy around the breathing frequency of 5 to 30 breathes per minute; and    (b) lower frequency local and systemic vasomotor effects on venous blood volume at frequencies from about 0.1 Hz down to below 0.01 Hz.
This signal can also be normalised in a similar manner to the pulsatile arterial signal described above.
As before, a normalising value can be produced by averaging a period of received signal, or passing the complete signal through a low-pass filter set to remove frequencies above, say, 0.05 Hz, so producing a signal value representing the overall effect of scattering and reflection, tissue absorption, non-pulsatile arterial blood absorption, non-time dependent venous blood absorption and averaged low frequency vasomotor time dependent venous blood absorption.
However, there are also applications where important physiological information is contained in the amplitude of the breathing signal frequency range of venous blood volume variation. In particular, consider an application where a significant amplitude difference in this frequency range, when comparing a pair of transducers in different anatomical locations, is indicative of a certain physiological condition. For example, the transducer pair may be located on contra-lateral parts of the body, such as the left and right upper arms, or the transducers may be mounted on left and right legs below the knee, where the physiological condition of interest may for instance be a lower limb DVT, either above or below the knee.
In such a case, the standard method of normalisation with the “dc” signal value described above introduces some problems. The frequency ranges of interest for vasomotor and breathing effects overlap with each other, so an ideal “dc” signal cannot be produced, containing the former with none of the latter, or vice versa. In addition accurate averaging of such low vasomotor frequencies requires long data records, which may not be available or consistent with a short diagnostic process.
The aim of such normalisation is to remove any effect on signals which does not affect each signal in the pair in the same way. Such differences may be in electronic hardware tolerances, sensor location or anatomy, as follows.
Hardware tolerances include:    (a) tolerances in overall circuit gain, including optical transmitter and receiver transduction ratios;    (b) tolerances in analogue filter frequency responses, particularly the low frequency high-pass “dc” level removing filter; and    (c) variations with time and temperature in these parameters, which are different between transducers.
Sensor location differences include:    (a) skin areas with differing amounts of hair;    (b) kin conditions on one leg which require the constant use of creams or moisturisers which include certain dermatological conditions and skin drying caused by long term use of anti-coagulants; and    (c) medical conditions on one leg which cause stretching of the skin at the transducer location such as oedema or the swelling associated with post phlebotic syndrome.
Anatomical differences between contra-lateral sites include:    (a) varying pigmentation;    (b) varying degrees of healed scar tissue from leg or pressure ulcers or other wounds; and    (c) variations in tissue attenuation caused by natural tolerances in tissue structures.
There may also be other problems which affect only one signal of a pair on a momentary basis, such as involuntary patient movement which brings the transducer into contact with other surfaces. Cabled transducers can also suffer from tension problems due to cable weight and stiffness and inevitable cable snags, as well as tension variations created by involuntary patient movement.
Note that in this latter case, for an asymmetric momentary problem, the averaging or filtering involved in producing a “dc” signal value will generate an asymmetric error which is impossible to remove. Conventionally, such an error can only be prevented by using a sophisticated pattern recognition algorithm, should one be possible, to identify the occurrence of a problem and remove use of this time slot of data on both signals.
In practice, in order to provide better received signal-to-noise ratios, IR optical transmitters are driven by high level current sources. In order to prevent thermal overload, and consequent thermal runaway and chip destruction, the drive is not continuous, but composed of narrow pulses, typically at a 1 to 3 kHz rate. The transmit pulse modulation signal is used to control a synchronous receiver.
To detect and compensate for any differential error between a plurality of transducer signals, including those listed above, an additional signal is introduced into the system.